The present invention relates generally to diagnostic imaging and, more particularly, to a method and apparatus of statistical recovery of pile-up events in a photon counting detector.
Typically, in computed tomography (CT) imaging systems, an x-ray source emits a fan-shaped beam toward a subject or object, such as a patient or a piece of luggage. Hereinafter, the terms “subject” and “object” shall include anything capable of being imaged. The beam, after being attenuated by the subject, impinges upon an array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is typically dependent upon the attenuation of the x-ray beam by the subject. Each detector element of the detector array produces a separate electrical signal indicative of the attenuated beam received by each detector element. The electrical signals are transmitted to a data processing system for analysis which ultimately produces an image.
Generally, the x-ray source and the detector array are rotated about a gantry within an imaging plane and around the subject. X-ray sources typically include x-ray tubes, which emit the x-ray beam at a focal point. X-ray detectors typically include a collimator for collimating x-ray beams received at the detector, a scintillator for converting x-rays to light energy adjacent the collimator, and photodiodes for receiving the light energy from the adjacent scintillator and producing electrical signals therefrom.
Typically, each scintillator of a scintillator array converts x-rays to light energy. Each scintillator discharges light energy to a photodiode adjacent thereto. Each photodiode detects the light energy and generates a corresponding electrical signal. The outputs of the photodiodes are then transmitted to the data processing system for image reconstruction.
Conventional CT imaging systems utilize detectors that convert radiographic energy into current signals that are integrated over a time period, then measured and ultimately digitized. A drawback of such detectors however is their inability to provide data or feedback as to the number and/or energy of photons detected. That is, conventional CT detectors have a scintillator component and photodiode component wherein the scintillator component illuminates upon reception of radiographic energy and the photodiode detects illumination of the scintillator component and provides an electrical signal as a function of the intensity of illumination.
While it is generally recognized that CT imaging would not be a viable diagnostic imaging tool without the advancements achieved with conventional CT detector design, a drawback of these detectors is their inability to provide energy discriminatory data or otherwise count the number and/or measure the energy of photons actually received by a given detector element or pixel. That is, the light emitted by the scintillator is a function of the number of x-rays impinged as well as the energy level of the x-rays. Under the charge integration operation mode, the photodiode is not capable of discriminating between the energy level or the photon count from the scintillation. For example, two scintillators may illuminate with equivalent intensity and, as such, provide equivalent output to their respective photodiodes. Yet, the number of x-rays received by each scintillator may be different as well as the x-rays intensity, but yield an equivalent light output.
Recent detector developments have included the design of an energy discriminating, direct conversion detector that can provide photon counting and/or energy discriminating feedback with high spatial resolution. In this regard, the detector can be caused to operate in an x-ray counting mode, an energy measurement mode of each x-ray event, or both. These energy discriminating, direct conversion detectors are capable of not only x-ray counting, but also providing a measurement of the energy level of each x-ray detected. While a number of materials may be used in the construction of a direct conversion energy discriminating detector, semiconductors have been shown to be one preferred material. Typical materials for such use includes Cadmium Zinc Telluride (CZT) or Cadmium Telluride (CdTe) having a plurality of pixilated anodes at attached thereto.
A drawback of direct conversion semiconductor detectors, however, is that these types of detectors cannot count at the x-ray photon fluxes typically encountered with conventional CT systems, e.g. at or above 106 counts per sec per millimeter squared. Saturation can occur at detector locations wherein small subject thickness is interposed between the detector and the radiographic energy source or x-ray tube. These saturated regions correspond to paths of low subject thickness near or outside the width of the subject projected onto the detector fan-arc. In many instances, the subject is more or less circular or elliptical in the effect on attenuation of the x-ray flux and subsequent incident intensity to the detector. In this case, the saturated regions represent two disjointed regions at extremes of the fan-arc. In other less typical, but not rare instances, saturation occurs at other locations and in more than two disjointed regions of the detector. In the case of an elliptical subject, the saturation at the edges of the fan-arc is reduced by the imposition of a bowtie filter between the subject and the x-ray source. Typically, the filter is constructed to match the shape of the subject in such a way as to equalize total attenuation, filter and subject, across the fan-arc. The flux incident to the detector is then relatively uniform across the fan-arc and does not result in saturation.
What can be problematic, however, is that the bowtie filter may not be optimal given that a subject population is significantly less than uniform and not exactly elliptical in shape. In such cases, it is possible for one or more disjointed regions of saturation to occur or conversely to over-filter the x-ray flux and create regions of very low flux. Low x-ray flux in the projection will ultimately contribute to noise in the reconstructed image of the subject.
“Pile-up” is a phenomenon that occurs when a source flux at the detector is so high that there is a non-negligible possibility that two or more X-ray photons deposit charge packets in a single pixel close enough in time so that their signals interfere with each other. Pile-up phenomenon are of two general types, which result in somewhat different effects. In the first type, the two or more events are separated by sufficient time so that they are recognized as distinct events, but the signals overlap so that the precision of the measurement of the energy of the later arriving x-ray or x-rays is degraded. This type of pile-up results in a degradation of the energy resolution of the system. In the second type of pile-up, the two or more events arrive close enough in time so that the system is not able to resolve them as distinct events. In such a case, these events are recognized as one single event having the sum of their energies and the events are shifted in the spectrum to higher energies. In addition, pile-up leads to a more or less pronounced depression of counts in high x-ray flux, resulting in detector quantum efficiency (DQE) loss.
Direct conversion detectors are also susceptible to a phenomenon called “polarization” where charge trapping inside the material changes the internal electric field, alters the detector count and energy response in an unpredictable way, and results in hysteresis where response is altered by previous exposure history.
This pile-up and polarization ultimately leads to detector saturation which occurs at relatively low x-ray flux level thresholds in direct conversion sensors. Above these thresholds, the detector response is not predictable and has degraded dose utilization that leads to loss of imaging information and results in noise and artifacts in X-ray projection and CT images. In particular, photon counting, direct conversion detectors saturate due to the intrinsic charge collection time (i.e. dead time) associated with each x-ray photon event. Saturation will occur due to pulse pile-up when x-ray photon absorption rate for each pixel is on the order of the inverse of this charge collection time.
Photon counting systems typically have one or more energy bins that are determined by a comparator that typically is part of the readout DAS. For a one-bin system, typically one energy threshold of the comparator is set to an energy value that is high enough such that there are few or no false noise counts, but low enough such that there is little loss of signal x-rays in the readout process. Such a system is subject to statistical error and bias due to the pile-up of multiple energy events, as described.
A system having many energy bins may be formed with multiple comparators in the readout DAS. Each comparator may be set to trigger for photons above a set level of energy that results in accumulation on a register of the number of photons above a corresponding x-ray energy level. The bin counts may be weighted and added together to form a system output having specific information content appropriate for an imaging system. However, like a one-bin system, a multiple-bin system is subject to degradation due to pile-up, resulting in DQE loss. The mean pile-up of bin counts may be corrected, but with a loss of statistical accuracy. The signal-to-noise ratio (SNR) may be used to assess the weighted sums for a system output. The DQE may be determined as 1/(1+N/N0), where N0=1/deadtime. The DQE may likewise be described as SNR2 (output)/SNR2 (input).
As an example, for a constant flux incident upon a non-paralyzable photon counting system with one comparator (i.e., one energy bin), the mean output counts M is related to the mean actual number of events N in the detector by the DQE as a multiplicative factor, which quantifies the statistical information loss as decreased SNR2. In other words, M=DQE*N. A Poisson distribution may be recovered if each output count is divided by this factor M, however, there will still be a loss of statistical information as quantified by the DQE. Similarly, for detectors having multiple energy bins, deconvolution techniques exist for correction of the mean spectral distortion created by pile-up. However, each energy bin likewise experiences a loss of SNR2.
Therefore, it would be desirable to design an apparatus and method improving statistical recovery of pile-up events through loss of DQE.